polyglycerol based hydrogels for the immobilization of catalytically
TRANSCRIPT
Polyglycerol Based Hydrogels for the
Immobilization of Catalytically Active Enzymes
and as Scaffolds for Cells
Inaugural-Dissertation
to obtain the academic degree
Doctor rerum naturalium (Dr. rer. nat.)
Submitted to the Department of Chemistry, Biology, and Pharmacy
of Freie Universität Berlin
by
Pradip Dey
From Howrah, India
August, 2015.
This Ph.D thesis was carried out under the supervision of Prof. Dr. Rainer Haag from
December, 2010 until August, 2015 at the Institute of Chemistry and Biochemistry of the
Freie Universität Berlin. The first three years of studies have been supported by Berlin
International Graduate School for Natural Sciences and Engineering (BIG NSE) and the final
years have been funded by Dahlem Research School (DRS).
1. Reviewer: Prof. Dr. Rainer Haag, Freie Universität Berlin
2. Reviewer: PD Dr. Kai Licha, Freie Universität Berlin
Day of Defense: 09.10.2015
To My Family
Acknowledgements
I want to express my sincere gratitude to my Doctoral supervisor Prof. Dr. Rainer Haag for
giving me the opportunity to be a member of his outstanding research group. I thank him for
the academic and scientific support he extended to me during this last five years. Also I want
to thank him for giving me the opportunity to work on such interesting projects and expose
me to all the aspects of research.
I would like to thank Dr. Kai Licha for his consent to be the second reviewer of my thesis and
taking part in subsequent defense procedure.
I would like to say thanks to all the co-operation partners I have been working with during my
Ph.D. I am particularly thankful to Prof. Ulla Wollenberger (University of Potsdam), Dr.
Gundula Schulze-Tanzil (Charité), Prof. Dr. Michael Gradzielski (Technical University of
Berlin), and Dr. Timothy Kennedy (McGill University, Canada) for fruitful collaboration and
also allowing me to work in their labs. This interdisciplinary thesis can never been successful
without the help from all the involved PhD students and post docs from several groups.
Special thanks to Tobias Schneider (Charité) for carrying out all the studies with
chondrocytes. I thank Dr. Miriam Adamovsky for introducing me to the electrochemistry and
also helping me with the measurements. I also thank Dr. Leonardo Chiappisi, Dr. Artavazd
Badalyan, and Dr. Radu-Cristian Mutihac for suggestions and measurements.
I would like to thank all the present and past group members of Haag group for their support
and providing a nice working environment. I thank Jutta Hass for taking care of my
scholarships and official documentation. I thank Dr. Pamela Winchester for proofreading my
manuscripts and thesis. I thank Dr. Wiebke Fischer for helping me always. I also
acknowledge Gaby Hertel, Katharina Goltsche, Cathleen Schlesener, Anja Stöshel, and
Andrea Schulz for taking care of chemical ordering, GPC measurements, and help in
measurements. I thank Dr. Florian Paulus and Daniel Stӧbner for the synthesis of dendritic
polyglycerol and MALDI TOF measurements.
I thank Dr. Florian Mummy and Dr. Dirk Steinhilber for helping me to establish in the lab. I
am grateful to Dirk for discussions and always he was my inspiration during PhD. I thank my
office and lab mates Dr. Maike Lukowiak, Dr. Emanuel Fleige, Dr. Sumati Bhatia, Dr. Juliane
Keilitz, Dr. Indah Nurita Kurniasih, Fatemeh Sheikhi Mehrabadi, Dr. Changzhu Wu, Karolina
Walker, Fang Du, Michael Unbehauen, and Pallavi Kiran for keeping friendly lab and office
environment. I acknowledge Maike’s efforts to have fun between the lab-mates such as lab
dinner, dancing in the lab, listening to songs, and she was always helpful and supportive. I
thank Dr. Sumati Bhatia to bear me in the lab, in the office, and also during lunch in last 3.5
yrs. I wish her success in all parts of life as she wants. Furthermore, I acknowledge her efforts
for proofreading my manuscripts and this PhD thesis. I thank Fang and Karolina especially for
all the discussions about polycaprolactone in last one year. After each discussion we came up
with new findings. I wish success in their PhDs. Also I thank Karolina for the translation of
the short summary of my thesis.
I specially acknowledge Dr. Dominic Gröger, Dr. Katharina Achazi, Markus Hellmund,
Sabine Reimann, and Mathias Dimde for their help in several occasions and friendship.
Next I thank Dr. Balu Thota for all the discussions and activities. I wish him success and
happy prosperous future. I acknowledge Dr. Shilpi Gupta, Dr. Shaswat Malhotra, Dr, Rahul
Tyagi, and Dr. Manoj Kumar Muthyala for social activities.
I acknowledge all the research facilities in the Institute of Chemistry and Biochemistry, Freie
Universität Berlin and would like to thank Dr. Andreas Schäfer, Dr. Andreas Springer, and
their colleagues for NMR and mass measurements.
During last five years I met with nice human beings from different countries. I thank Jun.
Prof. Dr. Marcelo Calderon, Dr. Chris Popeney, Dr. Mohiuddin Abdul Quadir, Prof. Dr.
Mohsen Adeli, and Mazdak Asadian Birjand for all the fun activities (such as go-cart, paint
ball) and friendship they have extended towards me in previous years. Without this it may
have been difficult to adjust in a foreign country.
I acknowledge the Berlin International Graduate School for Natural Sciences and Engineering
(BIG NSE) scholarship from UNICAT for the funding for three years. I think without that my
dream of doing PhD in Germany would have never been possible. I thank Dr. J P Lonjaret for
his guidance and support to deal with the bureaucracy of Germany especially in the initial
three months. I acknowledge all my batchmates of WS2010. Special thanks to Swantje, Jian,
Xinrui, Rashed, Arno, and Amandine for their friendship and help.
Five years is not a short time. Without family members is really hard to stay alone in a foreign
country. But I think I am lucky to have Subrata Da and Payel as apartment mate. I always had
homely feeling. I also specially acknowledge Dipti Da, Subhamoy Da, Anup Da, Amrita Di,
Arpan, Partha and Amitava for their friendship and social activities. I acknowledge Arpan and
Amitava for cooking nice foods and meaningless discussions every day.
I acknowledge all the teachers I came across during my previous studies. Everybody
contributed for the development of my personality and knowledge. I acknowledge Dr.
Achintya Kumar Sarkar, Dr. Chandan Saha, BCK sir, SG madam, Dr. Arnab Halder, and Dr.
Pradipta Ghosh for their inspiration for choosing this way of life.
I feel really lucky to have all the friends who helped me in different stages of my life. I
mention few names Rabi, Subhrajit, Chandan, Saikat, Ranjit, Tapas, Somenath, Ramesh,
Pritha, Ria, Pujarini, Subhrakanti, Soumen, Samik, Saona, Pradipta, Biplab Sarkar, Biplab
Dutta, Sumit, Subhadip (nato).
I acknowledge all my family members for their support and confidence on me. I feel lucky to
have my parents affections and blessings with me which help me to get all the success in my
life. I also acknowledge my elder brother’s support and encouragements for gaining this
degree and taking care of my responsibilities towards my family. I acknowledge Shilpi’s
patience for keeping faith on me that I can finish this PhD and also her encouragement, love,
and support. I hope I can spend more time with them, in future.
I should stop here. Please forgive me if I miss somebody.
Contents
1 Introduction……..…………………………………………………………………….1
1.1 Hydrogels………………………………………………………………………………1
1.1.1 Hydrogel Swelling.…………………………………………………………………..1
1.1.2 Dendritic Polyglycerol for Protein and Cell Encapsulation......................................2
1.2 Enzyme Based Biosensors……….…………………………………………………….5
1.2.1 Definition, General Information.…………………………………………………....5
1.2.2 Enzyme Based Electrochemical Biosensors.………………………………………...5
1.2.3 Enzyme Immobilization ……………………………………………………………..6
1.2.4 Polymeric Scaffolds Used for Enzyme Immobilization……………………………...8
1.2.5 Crosslinking Chemistry…………………………………………………………….10
1.3 Degenerative Joint Diseases…………………………………………………………..12
1.3.1 Physiopathology of Osteoarthritis…………………………………………………12
1.3.2 Treatments for Osteoarthritis……………………………………………………....14
1.3.3 Cartilage Tissue Engineering……………………………………………………...15
1.3.3.1 Scaffolds………………………………………………………………………16
1.3.3.2 Cell Sources…………………………………………………………………...17
1.3.4 Biocompatible Polymers Used for Cartilage Tissue Engineering…………………17
1.3.4.1 Natural Polymers……………………………………………………………...17
1.3.4.2 Synthetic Polymers……………………………………………………………21
1.3.5 Cross-linking chemistry……………………………………………………………22
1.3.5.1 Covalent Crosslinking………………………………………………………...22
1.3.6 Hydrogel Degradation……………………………………………………………..27
1.3.7 Requirements for Perfect Cartilage Tissue Engineering Scaffolds………………..29
1.4 Heparin Mimetic Synthetic Polymers…………….…………………………………..30
2 Scientific Goals………………………………………………………………………32
3 Manuscripts & Publications………...………………………………………………34
3.1 Dendritic Polyglycerol−Poly(ethylene glycol)-Based Polymer Networks for
Biosensing Application ………………………………………………………………34
3.2 Mimicking of Chondrocyte Microenvironment Using In Situ Forming
Dendritic Polyglycerol Sulfate Based Synthetic Polyanionic Hydrogels.....................62
3.3 Hydrolytically Degradable, Dendritic Polyglycerol Sulfate based Injectable Hydrogels
using Strain Promoted Azide-Alkyne Cycloaddition Reaction …………………….105
4 Summary & Conclusion…………………………………………………………...128
5 Outlook……...………………………………………………………………………130
6 Zusammenfassung…………………..……………………………………………...131
7 References…………………………………………………………………………..133
8 Appendix……………………………………………………………………………140
8.1 Publication & Manuscripts………...………………………………………………...140
8.2 Conference Contributions…………………….......…………………………………140
8.3 Curriculum-Vitae……………………………………………………………………142
i
List of Abbreviations
ACT autologous chondrocyte transplantation
BMP bone morphogenetic proteins
CS chondrotin sulfate
DF degree of functionalization
ECM extracellular matrix
dPG dendritic polyglycerol
dPGS dendritic polyglycerol sulfate
FGF fibroblast growth facor
GAG glycosaminoglycan
GOx glucose oxidase
HA hyaluronic acid
IL interleukin
IGF insulin-like growth factor
MMP matrix metalloproteinases
OA osteoarthritis
PAMAM poly(amido) amine
PaoABC periplasmatic aldehyde oxidoreductase
PEG poly(ethylene glycol)
PEG-DA poly(ethylene glycol) diacrylate
PEG-DGE poly(ethylene glycol) diglycidylether
PSS poly(4-styrenesulfonic acid)
PVA poly(vinyl alcohol)
PCL polycaprolactone
SPAAC strain promoted azide alkyne cycloaddition
TNF tumor necrosis factor
3D three-dimensional
Introduction
1 Introduction
1.1 Hydrogels
Hydrogels are water-swollen, three-dimensional (3D) polymer networks, which are classified
by several factors including the source of polymer and crosslinking being used and the
network’s nature and fate in the organism.[1-2]
Depending on the source of polymers used
hydrogels fall into one of three categories: natural, synthetic, or hybrid. According to the
nature of their crosslinking, they are either covalent or physical/supramolecular gels. The
nature of the network dictates whether it is homopolymer, copolymer, or interpenetrating.
Hydrogels can be classified as degradable and non-degradable by their fate in the organism.[2]
Lim et al. described the first hydrogel synthesis in 1960 when they copolymerized triethylene
glycol monomethacrylate with triethylene glycol dimethacrylate.[2]
Since then, there has been
an increasing urge to use the hydrogels for biomedical applications.[1-2]
1.1.1 Hydrogel Swelling
Hydrogel structures are mostly biocompatible due to their high water content. They swell
upon exposure to water because of the osmotic pressure that is created within hydrogels by
hydrophilic polymers, which creates a three-dimensional network. The swelling process can
be divided in three steps: (i) the water molecules diffuse through the hydrogel matrix, (ii) the
polymer chains are hydrated, and (iii) causes the polymer network to expand. The balance
between the thermodynamic force of mixing and the retractive force of the three-dimensional
(3D) network affects and controls the water content at the equilibrium swollen state.[1]
The
hydrophilicity of the polymer backbone, which contributes to the thermodynamic force of
mixing, is characterized by the interaction parameter (χ). The retractive force depends on the
number of crosslinks connecting the polymer chains into a 3D network.[3]
The network
expansion stops at equilibrium as these forces become balanced. The degree to which the
hydrogel swells depends on several factors like the osmotic pressure inside the hydrogel and
the crosslinking density.[4]
Altering either one of these factors increases or decreases the
degree of swelling. For example, osmotic pressure can change in a hydrogel with the
deprotonation of carboxylic acids due to a change in pH, whereas crosslinking density
changes with the degradation of the network.[4]
1
Introduction
1.1.2 Dendritic Polygylcerol for Protein and Cell Encapsulation
The adsorption of protein on implant surfaces is avoided by coating the surfaces with
hydrophilic polymers (protein resistant polymers).[5]
The main idea is to have minimum
interaction with protein and polymer (or with metal surface). Similarly in the development of
enzyme based biosensors (or enzyme based biofuel cells), the original enzymatic activity on
the electrode surface can be retained by avoiding or minimizing the interaction of enzymes
(protein) with the polymeric scaffolds. Our main assumption is that the minimum interaction
of proteins with the matrix can lead to a maximum stabilization of the protein. Dendritic
polyglycerol (dPG) is a branched hydrophilic, non-toxic, and anti-fouling polymer containing
multiple hydroxyl (OH) groups on the periphery/surface (Figure 1a).[5-7]
dPG is mainly
prepared by an anionic, ring-opening, multi-branching polymerization with narrow
polydispersity.[8]
This assumption has recently been proven (explored) in solution by
encapsulating pharmaceutically relevant proteins like asparaginase and lysozyme in dPG
based degradable nanogels (Figure 2). The activity of asparaginase retained even after release
from the nanogels after encapsulation.[9]
Thus, dPG have all the characteristics to become an
excellent matrix for enzyme (protein) encapsulation or immobilization because of its anti-
fouling properties.[10-11]
The presence of easily functionalizable hydroxyl groups makes them
good candidates for use as macromolecular crosslinkers for preparing multifunctional
polymeric network/hydrogels.[12]
The network or crosslinking density can easily be tuned by
varying the degree of functionalization (DF) of the dPG. By using dPG, the ligands that are
responsible for cell spreading and attachment, for example, RGD (Arg-Gly-Asp) derived from
fibronectin, can be readily incorporated in dPG using the functional available groups.
Figrure 1. (a) Structure of dPG, (b) Illustrates the antifouling properties of glycerols
(architectures - linear, hyperbranched) studied by immobilzing them on different surfaces and
2
Introduction
anchoring ligands used for immobilization; Adapted with permission from ref. [5]
. Copyright
2014 WILEY-VCH.
Figure 2. (a) Illustrates asparaginase encapsulation in dPG based nanogels by
nanoprecipitation, (b) activity of asparaginase after release from degradable nanogels,
Adapted with permission from ref. [9]
. Copyright 2012 Elsevier B.V.
In 2006 Hennink et al. reported the first dPG based hydrogel synthesis. He used photo-
initiated polymerization of methacrylate functionalized dPG (Figure 3).[13]
They embedded
multi-potent stromal cells (MSC) in the hybrid hydrogels of hyaluronic acid (HA) and
dPG.[14]
The hydrogels were formed by the photo/UV polymerization of methacrylated HA
and dPG. The cell viability was 75% and they were able to show the compatibility of dPG for
cell encapsulation.[14]
In 2011 the Haag group successfully encapsulated yeast cells in the
dPG-PEG based microgels by using redox initiated chain growth acrylate polymerization.[10]
They observed only 30% cell viability after gelation but could see 80% cell viability in dPG
solution without the initiation.[15]
As a result, they concluded that uncontrolled radical
formation led to cell death. Next Seiffert et al. applied a more cyto-compatible crosslinking
chemistry such as thiol-Michael addition to form cell laden dPG-PEG based microgels.[16]
This strategy increased cell viability to more than 90%. Recently a more bioorthogonal strain
promoted azide alkyne cycloaddition (SPAAC) reaction has been applied for the formation of
dPG-PEG based cell laden microgels in the Haag group.[17]
Fibroblast NIH3T3 cells were
encapsulated in these microgels and were viable in the dPG-PEG based microgels after
release even a week later. The degradability was introduced by forming acetal bonds between
azide functionalized benzaldehyde derivatives and 1,2-diol of dPG (present in the
periphery).[17]
From the above discussions, it can be concluded that the dPG has the potential
3
Introduction
to become an excellent scaffold for cell encapsulation and thus can be used for tissue
engineering or cell delivery to a target site.
Figure 3. Description of the progress in the development of cell laden dPG-based hydrogels
using different crosslinking approaches such as photo-initiated free radical polymerization,
redox initiated radical polymerizations, thiol-Michael addition. Adapted with permission from
ref. [10, 13, 16]
. Copyright 2006 and 2010 Elsevier Ltd. Copyright 2011 American Chemical
Society.
4
Introduction
1.2 Enzyme Based Biosensors
1.2.1 Definition, General Information
A biosensor is an analytical device which measures physicochemical changes occurring in a
biological recognition layer that has been attached to a solid transducer.[15, 18]
The biological
recognition system is usually a receptor protein, antibody, or enzyme.[18-19]
The information is
translated from the biochemical domain into a chemical or physical output signal using a
transducer with a specific sensitivity. Depending on the nature of the transduction signal,
biosensors are mainly three different types: optical, electrical, and mechanical (Figure 3).[15,
20]
Figure 4. (a) Schematic of a biosensor depending on the kind of physical change, which is
measured in the transducer surface. Adapted with permission from ref. [15]
. Copyright 2012
Royal Society of Chemistry, (b) A schematic of an enzyme based mediated electrochemical
biosensor. Adapted with permission from ref. [21]
. Copyright 2013, Royal Society of
Chemistry.
1.2.2 Enzyme Based Electrochemical Biosensors
An electrochemical device serves as a transduction element in electrochemical biosensors.
Operational simplicity, low production costs, and suitability for real time detection make the
electrochemical biosensor interesting for practical usage.[22-24]
The principle of enzyme based
amperometric (electrochemical) biosensors is schematically described in Figure 4. In brief,
enzymes or proteins are mainly immobilized/entrapped in a polymer scaffold which is
attached to an electrode surface.[25-26]
If enzymes are directly immobilized on the electrode,
they can be quickly denatured. The substrate diffuses into the matrix and is recognized by the
enzyme. For example the substrate can be oxidized and the enzyme is reduced. If the enzyme
5
Introduction
can directly interact with the electrode, the enzyme transfers one electron to the electrode. In
amperometry the potential remains constant and an electrical signal is detected.[18, 27]
Depending on this electron transfer process from enzyme to electrode, three different kinds of
biosensors are developed: first-generation biosensors (where oxygen mediates the electron
transfer from enzyme to electrode), second-generation biosensors (where soluble redox
molecules like ferricyanide, ferrocene, etc., shuttle the electron from enzyme to electrode),
and third generation biosensors (enzymes that directly communicate with the electrode).[23-24,
28-32]
1.2.3 Enzyme Immobilization
Enzyme based biosensors are interesting because of their superior selectivity and high affinity
of the enzymes towards substrates.[18, 30]
In general, enzymes are larger than the substrates to
which they bind. Molecular recognition is achieved by the well-known lock and key principle
between the respective receptor molecule and the analyte. Furthermore most enzymes are not
sufficiently stable at operational conditions. Therefore the immobilization of enzyme is a
crucial step for biosensor fabrication.[33]
There are five methods for immobilizing enzymes on
support materials: adsorption, covalent attachment, entrapment/encapsulation, and
crosslinking (Figure 5).[34-35]
Figure 5. Common immobilization methods involved in the fabrication of enzyme electrodes
for biosensor development. Adapted with permission from ref. [18]
. Copyright 2010 Royal
Society of Chemistry.
6
Introduction
Simple adsorption of enzymes on a support is a less complex method (Figure 5b).[27]
Mainly
electrostatic interactions like H-bonding interactions, van der Waal interactions, ionic forces,
and hydrophobic interactions are responsible for adsorption processes. In this method there is
no need for activation of enzyme or support; existing interactions are enough for the
immobilization.[18]
The immobilization is done by a simple and cheap procedure that consists
of mixing together the enzyme and support under suitable conditions of pH and ionic strength
for the period of incubation.[27]
Afterwards, the unbound biological component is removed by
washing the immobilized material extensively. The main disadvantage of this method is
leakage due to the reversible nature of interactions between the enzyme and support.[27]
Figure 6. Schematic description of covalent coupling of enzyme to the electrodes using
carbodiimide (a) and cyanogen bromide (b) as activating agent.
Covalent bonds are directly formed between the enzyme and solid/polymer support in
covalent binding (Figure 5c). Generally amino or carboxylic groups in the enzymes
(specifically in amino acid side chains) are used to immobilize the enzymes to the support by
amide or imine bonds. In most cases, two steps are involved: activation of the support and
enzyme attachment. Many reaction procedures are available for covalent coupling of an
enzyme and a support.[20, 27]
It is crucial to choose a method for maintaining the activity of the
redox site of the enzyme. During the coupling reaction, functional groups on the support (a
solid surface or polymer) should be activated first and then the enzyme (or protein) of interest
should be introduced. Glutaraldehyde or carbodiimide are mainly used for the covalent
7
Introduction
binding.[27]
The carbodiimide group activates the supports containing carboxylic acid (-
CO2H) and immobilizes the enzyme to the support via a peptide linkage (Figure 6a). Usually
the hydroxyl groups of polysaccharides are activated using cyanogen bromide (CNBr) to
couple the enzyme and support via an isourea linkage (Figure 6b).[20]
The main difference in the entrapment method compared to the adsorption and covalent
binding methods is that the enzymes are free in solution but their motion is restricted due to
crosslinking (Figure 5a).[36]
Generally enzymes can be entrapped in either of two ways:
behind a membrane or within a polymeric matrix. In the first case, a permeable membrane, a
thin film covering the electrode surface (detector), confines the enzyme solution
(suspension).[37]
In the other situation, enzymes are mixed with macromonomers to form
matrix.[38]
The most used technique for biosensor fabrications is entrapment in a polymeric
film (polypyrrole, Nafion) via casting or electropolymerization.[39-40]
The main disadvantage
of entrapment in a polymeric matrix is the possibility of enzyme leaching, which can be
prevented by controlling the gel porosity (or mesh size of networks). Gel porosity plays a
crucial role in controlling the mass transfer (free movements of substrates and products),
which is directly connected to the reaction kinetics and response time.
Crosslinking or co-crosslinking is mainly done by chemical or physical methods. This method
is support-free and involves connecting the enzymes (or biomolecules) to each other to form a
large, three-dimensional complex structure (Figure 5d). In the chemical method, a covalent
bond is formed between the enzymes using bi- or multifunctional reagents, such as
glutaraldehyde and toluene diisocyanate.[41]
To avoid the close proximity of the reactive site
(or catalytic site) other proteins like albumin and gelatin are sometimes used as good
molecular spacers. But for biosensor development, the co-crosslinking is rarely used alone as
a technique for immobilization, because of the possibility of enzyme denaturation and other
limitations arising from poor mechanical property and stability.
1.2.4 Polymeric Scaffolds Used for Enzyme based Biosensor Fabrication
Hydrogels (which contain cryo hydrogel, organogel, and grafting copolymer), sol-gel derived
organic-inorganic composites, and lipid membranes are used as membrane materials in the
field of enzyme based biosensors. For high stability and efficiency in a biosensor, the
biomolecules should be well separated to avoid self-aggregation and provide the same local
aqueous microenvironment as in the biological medium. In that sense, hydrogels which form a
crosslinked polymeric network are the most suitable scaffolds for such applications, because
8
Introduction
they can imbibe a lot of water and provide the proper protective microenvironment for
enzymes on the electrode surface. Polymers with several hydroxyl groups are interesting for
such applications. Polyvinyl alcohol (PVA) is a good matrix for immobilized enzymes
because of its good biocompatibility, chemical stability, and inertness to microbial
degradation.[42-43]
PVA is a linear commercially available polymer containing many hydroxyl groups, which is
synthesized from hydrolyzing corresponding polyvinyl acetates. Generally PVA films are
formed from photo-polymerization of corresponding macromonomers.[44-45]
The hydrolysis of
polyvinyl acetate never completes and makes it water insoluble.[44]
After PVA, the most readily available polymer used to form matrix for enzyme-based
biosensors is polyethylene glycol (PEG) which is well known to prevent nonspecific binding
and electrode fouling.[46]
PEG is mainly prepared by anionic ring-opening polymerization of
ethylene oxide. PEGs can be monofunctional, bifunctional, or a star PEG polymer. As
bifunctional PEG contain two hydroxy groups at the both ends, it is really easy to introduce
different functional groups like amine, thiol, vinyl sulfone, acrylate, methacrylate. Mainly
PEG diacrylate (PEG-DA) is used to crosslink the macromonomers by applying UV in the
presence of radical initiators.[47]
Dendrimers have been applied in drug delivery, energy harvesting, ion sensing, catalysis, and
information storage due to the presence of multivalent/polyvalent functional groups on the
periphery and the spherical nature of the particles.[48]
Among the various dendrimers,
poly(amidoamine) (PAMAM) dendrimers are used to develop enzyme based biosensors.[48-51]
Glucose oxidase (GOx) electrodes are mainly constructed using the PAMAM dendrimers as a
bio-conjugating reagent.[39]
PAMAM dendrimer and GOx are assembled layer by layer on the
electrode surface by forming imine bonds between the amino group from PAMAM and
aldehyde groups from the oxidized GOx.[39]
Reagent-less biosensors have also been
developed by introducing ferrocene aldehydes[52]
and Au-nano particles[53]
during the
formation of layer by layer assembly of PAMAM dendrimer and GOx. In a similar concept,
Pt DENs were synthesized by complexing G6-NH2 PAMAM dendrimers with Pt and further
immobilizing them on glassy carbon electrodes (GCEs) via an electrooxidative coupling of
the terminal amine groups of dendrimers to the carbon surfaces. Biotin-streptavidin
interactions have also been used to immobilize the GOx to the Pt DENs electrode.
9
Introduction
1.2.5. Crosslinking chemistry
Crosslinking chemistry plays a key role in controlling and retaining the activity of
encapsulated or immobilized enzymes. The crosslinkable groups (chemistry) should be
orthogonal to the functional groups present in the active center of enzymes or the groups
participating in biocatalysis. On this basis several crosslinking chemistries have been used to
prepare polymer films and enzyme based electrodes. Imine bond formations and
glutaraldehyde crosslinking are attractive due availability of amino groups on the enzymes.
Usually oxidized GOx are coupled to cystamine functionalized gold electrodes by an imine
bond, which is then reduced to secondary amine by sodium cyanoborohydride (NaBCNH3)
(Figure 7).[39]
Afterwards, the PAMAM dendrimers are immobilized to modified electrodes
for a layer-by-layer assembly on the electrode surface. By repeating the steps (Figure 7)
multilayered GOx/dendrimer network can be built on the gold electrode using the reductive
amination.[39, 52]
In a similar fashion, glutaraldehyde has been used as a crosslinking agent to
construct the enzyme electrode by utilizing the amino groups of the enzymes.[53]
Figure 7. Step by step multilayer film formation of GOx-PAMAM on gold electrode using
reductive amination. Adapted with permission from ref. [39]
. Copyright 2000 American
Chemical Society.
UV and photo initiated radical polymerizations are interesting because of their high reactivity
and easily available functional groups.[45, 47]
The ferrocene conjugated redox polymer
containing some benzophenone units is deposited on the surface of glassy carbon electrodes
by spin coating or drop casting. Afterwards a deposited polymer layer or film is irradiated by
UV to crosslink them using the photochemical reaction between benzophenone and adjacent
10
Introduction
C-H groups. The biocatalyst or active enzyme can be entrapped on the electrode surface by
performing the crosslinking in the presence of a biocatalyst. A redox polymer poly[4-
vinylpyridine Os(bipyridine)2Cl]-co-ethylamine is first immobilized on a 11-
mercaptoundecanoic acid (MUA) modified gold electrode by utilizing the ionic interaction.
Then the mixture of GOx and the solution of PEG-DA is drop casted on the redox polymer
modified surface. By irradiating the polymer film with UV, GOx is entrapped in the
biocompatible PEG-DA hydrogels. Above all, the electron exchange between the redox
polymer and the GOx entrapped in the hydrogel has also been observed in this situation.
Epoxide ring-opening reactions belong to the one of the click reactions described by Sharpless
et al. These are the most frequently used reactions to prepare enzyme entrapped redox
polymer films. Mainly amino groups react with the epoxides to initiate the crosslinking.
Heller et al. were the first ones to demonstrate the formation of a redox hydrogel film
containing GOx on a glassy carbon electrode by crosslinking osmium coordinatively bound to
poly(N-vinylimidazole) polymer and PEG diglycidyl ether (PEG-DGE).[54]
In this case, PEG
DGE is used as the crosslinker and the 3D network mediates the electron transfer to the
electrode via osmium- poly(N-vinylimidazole) complex.
From the above discussion it can be concluded that GOx has mainly been applied for the
development of sensors. But ongoing interests are focused on the development of biosensors
or biofuel cells using more robust and efficient enzymes. Nowadays, retention of long term
enzymatic activity on the electrode surface is a major challenge. So, there is an urgent need
for the design of suitable 3D scaffolds using easy and fast orthogonal crosslinking chemistries
to immobilize these efficient enzymes on the electrode surface.
11
Introduction
1.3 Degenerative Joint Diseases
Degenerative joint disease or osteoarthritis (OA) is an inflammation of the joints and
surrounding tissues.[55-56]
Overuse of joints, aging, or simply wear and tear is mainly
contributed towards degenerative joint diseases.[57]
In developed countries, OA is a major
concern and for example in the United States of America, persons with age older than 60
years need medical consultations primarily due to OA.[55]
OA is characterized by decrease in
articular cartilage (AC) thickness, subchondral bone sclerosis (bone thickening), formation of
osteophytes (bone outgrowth on the joint margin), and modification of the synovial fluid
composition.[56, 58]
Most commonly sites affected by OA are knees, hips, fingers, and the
lumbar and cervical spine.[59]
Unfortunately, AC has a limited capacity for self-renewal.
Osteoarthritic changes are usually diagnosed in an advanced stage due to difficulty in early
diagnosis. So, it is necessary to understand the underlying molecular mechanisms of cartilage
destruction in OA to develop and improve diagnostic and therapeutic approaches.[55]
1.3.1 Physiopathology of Osteoarthritis
Various connective tissues including AC, synovial membrane, subchondral bone, ligaments
and sometimes menisci are the parts of the diarthrodial joint as shown in Figure 8.[57]
The
joint function and performance originate from these complex structures. The chemical
composition of AC helps to execute the repetitive loading cycles. And the physical structure
contributes towards the essential frictionless motion. This unique viscoelastic and
compressive properties of AC originates from the extracellular matrix (ECM) of
chondrocytes.[60]
AC ECM is mainly composed of collagen type II and the large proteoglycan
aggrecan.[61-62]
The transportation of nutrients and cellular repair components to the
chondrocytes are mainly mediated by diffusion from the synovial fluid. Progressive loss of
the structure and functionality of AC occurs during the progression of OA due to an
imbalance between anabolic and catabolic processes in the tissue.[55]
Usually a complex
interplay of genetic, environmental, metabolic, and biochemical factors govern the AC
degradation in OA, although the correct mechanism is still unclear. Loss of collagen and
proteoglycans from the matrix was observed during the OA progression which may be due to
the result of over-expression of matrix degrading enzymes, such as matrix metalloproteinases
– MMPs).[55, 62-66]
It is also observed that in response chondrocytes start to produce more
matrix molecules such as collagen. As the disease progresses cartilage degradation overtakes
these repair attempts and leads to loss of cartilage. In OA cartilage the synovium and the
12
Introduction
chondrocytes produces the cytokines and growth factors, part of which are involved in
physiological cartilage development.
Figure 8. Simplified structure of a diarthrodial joint, adapted with permission from ref. [36]
.
Copyright 2012 Osteoarthritis Research Society International. Published by Elsevier Ltd.
Inset: structure of articular cartilage – extracellular matrix of chondrocytes. Reproduced with
permission from ref. [33]
. Copyright 2007 Nature Publishing Group.
MMP synthesis is increased on the catabolic side and the inflammatory cytokines like
interleukin (IL)-1, IL-17, and IL-18, and tumor necrosis factor (TNF)-α decreases the
synthesis of inhibitor of MMPs and extracellular matrix components.[56, 63]
On the other hand,
anabolic factors like insulin-like growth factor (IGF)-1, transforming growth factor (TGF)-β,
fibroblast growth factors (FGFs), and bone morphogenetic proteins (BMPs) stimulate the
synthesis of ECM components (Figure 9).[63]
The pathophysiological changes during OA
progression took place in three overlapping phases: first the matrix–network starts to slowly
degrade on a molecular level. This leads to an increase in water content and decrease in the
size of matrix polymers.[67]
The cartilage stiffness is reduced by damage to the collagen
network. In the second phase, metabolic activity and proliferation of chondrocytes are
enhanced to compensate the damage, i.e., cells synthesize new matrix molecules. This phase
lasts for several years. In phase three, cartilage loses all its tissue because the chondrocytes
are not able to keep up their repair activity.[61]
13
Introduction
Figure 9. Description of the catabolic and anabolic factors involved in the OA progression.
Reproduced with permission from ref. [55]
. Copyright 2009 Elsevier Ltd.
1.3.2 Treatments for Osteoarthritis
Until now different treatment approaches have been available for the treatment of OA, which
includes non-steroidal drugs, injectable hydrogels (for example Synvisc GF 20, acts as a
lubricating fluid), autologous chondrocyte transplantation (ACT), and joint replacement.
Acetaminophen, non-steroidal anti-inflammatory drugs (NSAIDs), cyclooxygenase-2 (Cox-2)
inhibitors, glucocorticoids, and opioids are mainly used for pain relief.[55]
But long-term use
of NSAIDs lead to adverse affects.[59]
ACT can be described in three steps: first chondrocytes
are isolated from the patient’s donor tissue, and expanded in vitro monolayer culture.[56]
Cells
are then transplanted to the defect site under a periosteal flap in the hope that it will regenerate
the cartilage tissue once again. ACT has been used clinically to repair articular cartilage
defects.[55]
But it is really hard to obtain high cell numbers from mature cartilage tissue using
biopsy. Also microfracture, mosaicplasty, ACT, and osteochondral allograft transplantation
are popular treatments nowadays that are available for AC repair.[68]
These techniques
discussed above have alleviated long-term knee pain but they have some limitations on their
own like the AC tissue produced applying these techniques is composed of collagen type I
(characteristically matches with fibro cartilage).[69]
This tissue is mechanically and
biochemically inferior to the original hylane cartilage mainly present in the cartilage.[69]
Also
the periosteal flap or the membrane used to keep the cells at defect site is not totally
impervious and sometimes leads to hypertrophy or uncontrolled calcification.[55, 68]
So, the
ultimate treatment available is joint replacement. Symptomatic relief can be effectively
achieved using the current surgical approaches, although they are not a complete and
14
Introduction
restorable solution. Donor side morbidity, formation of fibro cartilage, and lack of functional
integration with the host cartilage are their main limitations. To overcome all the problems
associated with OA treatments, three-dimensional scaffolds have been used to transfer and
maintain the cells at the recipient site.[70-73]
1.3.3 Cartilage Tissue engineering
The principles and methods of engineering and life sciences are applied in tissue engineering
for the development of biological substitutes to restore, maintain, or improve tissue
function.[74]
Mainly appropriate cells are seeded in a biocompatible scaffold.[56, 75]
Cell
differentiation and maturation can be induced by incorporating some signaling molecules
(morphogens) in the biomaterial scaffold.[55, 76]
There are two tissue-engineering approaches:
(1) functional tissue is generated in vitro by seeding the cells into a scaffold and the construct
is then implanted into the joint. (2) Cells are seeded in the scaffold, implanted, and allowed to
mature in vivo.[70]
Cartilage repair applying tissue engineering has gained a lot of attention
due to its poor capability for endogenous repair.[71]
Figure 10 describes the principle of
cartilage tissue engineering schematically.
Figure 10. Principle of cartilage tissue engineering. Reproduced with permission from ref. [55,
71]. Copyright 2009 Elsevier Ltd.
15
Introduction
1.3.3.1 Scaffolds
A 3D environment is required to mimic in vivo conditions for the production of cartilaginous
tissue, which can be achieved using the scaffolds. Cellular behavior is mainly controlled by
varying the scaffold architecture.[77]
So, delivery of cells in cartilage regeneration has been
done using numerous scaffolding materials.[70]
Hydrogels, sponges, and fibrous meshes
prepared using polymeric materials have mainly been applied for cartilage regeneration. In
situ forming and shear thinning hydrogels can be used as injectable scaffolds to easily fill
defects of any size and shape.[70, 78-79]
They can also be implanted in a minimally invasive
manner.[76]
Generally hydrogels contain lots of water which helps to transport nutrients and
waste.[80]
Cells can be mixed with the macromonomers solutions before the hydrogel
formation and therefore can be entrapped inside the polymer networks.[81]
Finally this will
lead to a homogenous suspension of the cells in a 3D environment. Encapsulated cells
typically retain a rounded morphology due to the absence of attachments around them, which
may induce a chondrocytic phenotype. The mechanical properties of hydrogels can be tuned
by varying the crosslinking density, which also can affect cell viability. So the major
drawback from using hydrogels remains their limited mechanical properties.
Sponges are interesting scaffolds for cartilage regeneration. Pore size, porosity, and
interconnectivity control the properties of sponge. Cell adhesion is dictated by the porosity.[80,
82] On the other hand, cell infiltration and migration, matrix deposition and distribution, and
nutrient and waste exchange are affected by pore size and the interconnectivity affect.[83]
Until
now sponges have been prepared by several methods such as porogen leaching, freeze-drying,
and gas foaming.[71]
These manufacturing methods affect the scaffold’s architecture, which in
turn controls tissue formation[82]
and can be used to encapsulate growth factors (GFs).[84]
Meshes are also applied for cartilage tissue engineering material because that they have high
void volumes and high surface areas. Meshes can be divided in two categories, woven and
non-woven fibers. Cell behavior is mainly controlled by variations in void volume and fiber
diameter and directionality. Woven meshes have better mechanical strength and they can be
prepared with a wide range of porosity. Cells can be seeded into these prefabricated scaffolds
and implanted in vivo for tissue regeneration. Complete integration to the cartilage cannot be
achieved with the prefabricated scaffolds.[71, 83]
This is the major drawback using meshes.
16
Introduction
1.3.3.2 Cell Sources
Chondrocytes, fibroblasts, stems cells, and genetically modified cells are being used for
cartilage tissue engineering.[70]
Chondrocytes derive from different places (articular, nasal,
and costal). The limited availability of chondrocytes makes it quite difficult to obtain enough
cells to fill a clinically relevant defect. Since 5-10% of cartilage tissue is composed of
chondrocytes, they have to expand in vitro for sufficient cell numbers. But the main problem
is that chondrocyte loses its phenotype stability during expansion in a monolayer culture.[69]
This reduction of phenotypic stability is termed dedifferentiation. Type II collagen expression
is decreased and type I collagen expression is increased during dedifferentiation.[69]
The cells
also transform morphologically from their normally rounded shape to the typical fiber like
shape of fibroblasts. Sometimes this process is reversible if the cells are cultured in a three-
dimensional environment. For cartilage tissue engineering, mesenchymal stem cells (MSCs)
can be considered an alternative source of reparative cells due to their chondrogenic
potential.[85]
For the most part, MSCs are separated from bone marrow and they have the
capacity to proliferate while retaining both their multipotency and ability to differentiate in
different lineages (like osteoblasts, chondrocytes, adipocytes, cardiomyocytes, and so on).[85]
1.3.4 Biocompatible Polymers Used for Cartilage Tissue Engineering
Both natural as well as synthetic polymers have been applied as a matrix or scaffold for
cartilage tissue engineering (Figure 9, and 10).[70-71]
Natural polymers are very interesting
because they can interact with cells via cell surface receptors, but this interaction can
stimulate an immune system response.[71]
Natural polymers can easily be degraded in
presence of the host enzymes. In addition, the mechanical property of natural polymer may be
inferior, whereas the chemical and degradation characteristics of synthetic polymers can be
controlled and tuned depending on the circumstances.[71]
Normally synthetic polymers do not
have direct cell scaffold interactions but they can be synthetically incorporated for better cell
adhesion, signaling, and directed degradation.[86]
The following is a detailed discussion about
the polymers used for cartilage repair including their advantages and disadvantages.
1.3.4.1 Natural Polymers
Natural polymers like alginate,[87-89]
agarose,[90-91],
[92] fibrin,[93]
hyaluronic acid (HA),[94-100]
collagen,[101-102]
gelatin,[103]
chitosan,[98, 104-105]
chondroitin sulfate,[98, 103, 106]
heparin,[107]
and
cellulose[83]
have been applied as bioactive scaffolds for cartilage tissue engineering.[70]
17
Introduction
However, the presence of endotoxins and their fast enzymatic degradation can be the limiting
factor to use natural polymers as cartilage tissue engineering scaffolds.
1.3.4.1.1 Chitosan Based Hydrogels
Partially deacetylated chitin is called chitosan, which is found in the exoskeletons of
arthropods and is a liquid at room temperature and a gel at physiological temperatures. N-
acetylglucosamine groups are randomly located in chitosan (see Figure 11a).[70]
Chitosan can
form an insoluble ionic or polyelectrolyte complex with various water-soluble anionic
polymers due to its cationic nature and presence of high charge density. The mechanical
properties of chitosan can be improved ionically or by covalent crosslinking. Normal
chondrocyte phenotypes can be kept in 2D and 3D cell cultures using chitosan and chitosan
hybrid hydrogels as a support.[108-109]
Figure 11. Structures of natural polymers such as (a) chitosan, (b) alginate, (c) agarose, (d)
hyaluronic acid, (e) chondrotin sulfate, (f) heparin are shown used as a scaffold for cartilage
tissue engineering. Adapted with permission from ref. [70]
. Copyright 2011 American
Chemical Society.
1.3.4.1.2 Alginate Based Hydrogels
Alginate is a copolymer of (1,4)-linked β-D-mannuronic acid and R-L-guluronic acid (Figure
11b). Alginate is negatively charged due to the presence of carboxylic acids. Bivalent cations
18
Introduction
like magnesium (Mg2+
), calcium (Ca2+
), and barium (Ba2+
) can be used as a crosslinking agent
to form alginate gels. The advantages of alginate gels are their easy preparation, favorable
cellular response, and low cost. These properties make them attractive candidates for
developing tissue-engineering constructs.[87]
A clear difference is observed in terms of
attachment, morphology, and chondrogenesis when articular chondrocytes are seeded in RGD
functionalized alginate gels and in functionalized gels without RGD.[71, 110]
A more flattened
morphology with stress fibers is observed upon increased crosslinking density and substrate
stiffness.[111]
The main limitations of alginate gels are their low mechanical properties and
slow degradation rates, although successfully chondrogenesis can be studied with alginate
gels in vitro.
1.3.4.1.3. Agarose Based Hydrogels
The repeating unit present in agarose, which is derived from Asian seaweeds, is agarobiose
(Figure 11c). Galactose and 3,6-anhydrogalactose are alternating units present in agarobiose.
Agarose gel can be formed by cooling down a homogeneous solution of agarose from 99 to
35 °C. Gelation occurs due to the transformation of coil to helix transitions of agarose. The
mechanical forces applied on the agarose hydrogels can be transmitted to cells during
compression, which make them suitable for studying the chondrocyte response to
deformational loading.[91-92, 112-113]
1.3.4.1.4 Hyaluronic Acid Based Hydrogels
Hyaluronan or hyaluronic acid (HA) is a glycosaminoglycan (GAG) that is natively present in
cartilage.[94]
The major component of synovial fluid is HA, which is comprised of D-
glucuronic acid and D-N-acetyl glucosamine units. They are linked together via alternating β-
1,4 and β-1,3 glycosidic bonds (Figure 11 d). HA is advantageous in cartilage tissue
engineering because it inhibits fibronectin fragment-mediated chondrocytic chondrolysis,[114]
prostaglandin synthesis, proteoglycan release, and degradation and that it is anti-
inflammatory.[70, 94, 115]
In addition, proteoglycans form aggregate with other GAGs, which is
mainly mediated by the HA present in cartilage. It can also be degraded by enzymes,
hyaluronidase, and free radicals. HA is the best candidate for cartilage tissue engineering
because it has free OH, COOH, and N-acetyl groups. For photopolymerization, several
methacrylate and acrylate groups can be introduced via esterification.[95, 100]
The poor
mechanical strength of HA based hydrogels is a drawback in cartilage tissue engineering. A
wide range of mechanical properties, however, can be achieved by varying the molecular
19
Introduction
weight and concentration of modified HA macromer.[116]
Increasing the macromer
concentration increases the compressive moduli and degradation time but this also causes the
swelling ratio and cell viability to decrease.[96]
HA based hydrogels have shown promise in
cartilage tissue engineering[99]
but availability and variability of the material properties, and
quality in terms of possible pathogen contamination still remains a major concern.
1.3.4.1.5 Chondrotin Sulfate Based Hydrogels
N-acetylgalactosamine and glucuronic acid are alternating sugars present in chondroitin
sulfate (CS), which is a sulfated GAG (Figure 11e). CS can prevent the prevalence of OA. It
is well known that CS has the capability to enhance or stimulate the metabolic response of the
tissue both in vitro and in vivo and that it has anti-inflammatory properties.[70]
CS can prevent
proteoglycan (PG) degradation in the osteoarthritis model in rabbits.[117]
CS can be degraded
by chondroitinase which is secreted by cells.[70]
CS based hydrogels can also be synthesized
from photopolymerization of the corresponding methacrylate macromer (methacrylation of
hydroxyl groups of CS). This CS based hydrogel inhibits the biosynthetic activity of
chondrocytes. The only limitation to using CS based gels is that CS is highly negatively
charged and attracts free cations from the medium, which results in an increase in osmotic
pressure within the hydrogels. This makes the CS based hydrogels unsuccessful for use in
cartilage tissue engineering. But the introduction of PEG in CS based hydrogels enhances the
chondrogenic gene expressions and cartilage matrix productions.[118-119]
As a result, it is better
to use hybrid hydrogels of CS (composed of synthetic or other biopolymers) for cartilage
tissue engineering than pure CS hydrogels.[120-121]
1.3.4.1.6 Collagen/Gelatin Based Hydrogels
Collagen is a biomacromolecule that is abundant in cartilage tissue[122]
and has all the
interesting characteristics for use as a carrier material for cartilage tissue engineering.[122], [102]
Chondrocyte can interact with type I and type II collagen scaffolds via the integrins present in
the cells. Collagen type II supports chondrogenesis of MSCs and enhances the effect of TGF
β1.[101, 113, 123-124]
Gelatin is a biopolymer derived from collagen. The only drawback of a
collagen scaffold is that is can easily degrade and thus lose the mechanical properties before
healing is complete.
20
Introduction
1.3.4.1.7 Heparin Based Hydrogels
Heparin is a negatively charged highly sulfated polysaccharide that is used as an
anticoagulant. Heparin is a linear glycosaminoglycan (GAG) and a heterogeneous mixture of
1,4-linked uronic acids (D-glucuronic, L-iduronic, or L-2- sulfated iduronic) and glucosamine
residues (D-N-acetyl glucosamine and D-di-N-6-sulfate glucosamine) (Figure 11f). Heparin
can interact with bioactive proteins that are associated with cell adhesion, proliferation, and
differentiation. Hydrogels containing heparin are interesting for biomedical applications such
as a controlled release of growth factors and cartilage regeneration.[107, 125-128]
The tissue
regeneration potential of heparin based hydrogels has also been examined. Heparin based
hydrogels are formed by incorporating thiol groups or tyramine groups, respectively, to linear
heparin chains for crosslinking with tetra-arm star PEG acrylate or tyramine conjugated
dextrans.[107]
1.3.4.2 Synthetic Polymers
Synthetic polymers like poly (α-hydroxy esters),[129-131]
polyethylene glycol (PEG),[123, 132-134]
polyvinyl alcohol (PVA),[120, 135-136]
poly(NiPAAm),[137-138]
poly(propylene fumarates),[139-141]
and polyurethanes[142]
have been applied as scaffold materials for cartilage tissue engineering
(Figure 12). Among them, PEG has been mostly used because of its inertness and
biocompatibility.[143]
Crosslinked PEG hydrogels support chondrogenesis. The incorporation
bioactive peptides, hydrolysable blocks like lactic acid,[132]
caprolactone,[144]
etc., in general
lead to better tissue formation and cell proliferation. In addition, PEG has been incorporated
in natural polymers as crosslinker for preparing hybrid hydrogels to improve their mechanical
properties and degradability.
21
Introduction
Figure 12. Synthetic polymers used as cartilage tissue engineering scaffolds
1.3.5 Crosslinking Chemistry
Crosslinking chemistry plays an important role in keeping better cell viability during the
encapsulation process in the hydrogels. In addition, the scaffolds play a crucial role in
providing enough mechanical stiffness to cartilage for in vivo applications. Hydrogels are
formed by covalent crosslinking and supramolecular or ionic interactions for cartilage tissue
engineering. Supramolecular or ionic crosslinking, however, leads to hydrogels with weak
mechanical properties compared to covalent crosslinking. Thus in the next sections, covalent
cross-linking chemistry will be discussed in the context of cartilage tissue engineering.
1.3.5.1 Covalent Crosslinking
The main goal of cartilage tissue engineering is to encapsulate the cells during the gelation
(i.e. crosslinking process). The crosslinking chemistry should be bioorthogonal to have a
minimum effect on the cell viability. Bioorthogonal reactions do not interfere with the
biological systems. Other requirements are that the reaction rate should be fast enough,
selective, and high yielding.
22
Introduction
1.3.5.1.1. Chain Growth Radical Crosslinking
The most frequently used crosslinking strategy for the encapsulation of cells in the hydrogels
is radical crosslinking due its fast reaction kinetics. Usually crosslinking can be initiated by
light, temperature, or redox conditions. The other advantage of radical reactions is the
introduction of reactive groups to macromonomers (functional polymers) for crosslinking is
straightforward. Reactive acrylate, methacrylate, or acrylamide groups can be introduced by
ester or amide bonds. Another advantage is visible light can be used to initiate the
crosslinking using the initiators (activated by visible light), so it can be applied for an in vivo
delivery of gel-cell construct.[145]
Figure 13. (a) Photoinitiated chain growth radical crosslinking of PEG-diacrylate (PEG-DA).
(b) photo-initiated thiol-ene step growth polymerization of star-tetra arm PEG norbornene and
PEG dithiol. Confocal laser scanning microscopy (CLSM) image of intracellular reactive
oxygen species (ROS) generated following chondrocyte encapsulation during chain growth
radical crosslinking of PEG-DA (c) and step growth thiol-ene reaction of star tetra arm PEG
norbornene and PEG-SH (d) (red stained with carboxy-H2DFFDA (carboxy-2,7-
difluorodihydrofluorescein diacetate). CLSM image of encapsulated chondrocytes after 24 h
in chain growth acrylate hydrogels (e), and step growth thiol-ene hydrogels (f) (green - live
cells stained with Calcein AM, red – dead cells stained with Ethidium bromide). Reproduced
from ref. [146]
. Copyright 2013 Elsevier Ltd.
But the major concern for using radical crosslinking is the formation of intracellular reactive
oxygen species (ROSs) during photo crosslinking, which may affect the cell viability due to
increase in tendency of DNA damage[145]
and tissue formation.[99, 147]
Recently, Bryant et al.
23
Introduction
compared the influence of reactive groups like acrylate and thiol-norbonene crosslinking on
chondrocyte encapsulation in terms of tissue formation (Figure 13).[146]
They have shown that
acrylate crosslinking lead to an encapsulation environment with an elevated intracellular ROS
compared to a thiol-norbornene system. Overall, the chondrocyte encapsulation in acrylate
system leads to a neo-tissue formation that resembles hypertrophic cartilage. Whereas the
neo-tissue is formed from chondrocyte encapsulation in a thiol-norbornene system, has the
properties resembles to hyaline cartilage after several weeks cell culture.[146]
In addition, a
major concern is that the free-radical reactions are highly exothermic. The light intensity
decreases in deep hydrogel which may lead to inhomogeneous crosslinking density.
In conclusion, radical crosslinking has been applied for cartilage tissue engineering due to the
availability of macromonomers and an easy synthesis but there is a need to use more bio-
orthogonal crosslinking approaches.
1.3.5.1.2 Thiol-Michael Addition Reaction
In a thiol-Michael addition reaction, thiol nucleophile usually adds to a conjugated double
bond (unsaturated ketone or amide, Michael acceptor), an acrylate, maleimide, or a vinyl
sulfone (Figure 14a).[148]
Under physiological conditions the reactions can be performed in an
aqueous medium, which make them suitable for the encapsulation of cells in hydrogels.
Peptide hydrogels are mainly prepared by thiol-Michael addition as thiol is one of the reactive
groups in this reaction.[149]
Thiols are readily available in peptides. The other reactive group
can be easily incorporated to polymer. HA-PEG and Dextran-PEG based hybrid injectable/in
situ forming hydrogels were prepared using thiol Michael addition for cartilage tissue
engineering.[100, 150]
HA is functionalized with thiol by coupling cystamine dihydrochloride to
the carboxylic acid group of HA, whereas cystamine is coupled to hydroxyl groups of dextran
by activating them using p-nitrophenyl chloroformate. Tetra-arm star PEG vinyl sulfone
(PEG-4VS) and PEG-acrylate (PEG-4-acr) are used as crosslinkers.
1.3.5.1.3 Thiol-ene Click Reaction
A thiol-ene reaction proceeds through a step growth polymerization. This reaction is mainly
performed between a thiol and a norbornene derivative in the presence of photo initiator and
light (365 nm) (Figure 14b). This reaction offers a more controlled radical formation
compared to the free radical polymerization.[146]
Brayant et al. applied this polymerization for
the encapsulation of chondrocytes, which led to better tissue formation compared to hydrogels
formed from a free radical polymerization (Figure 12b).[146]
As thiol is one of the reactive
24
Introduction
groups in the thiol-ene click reactions, peptides can easily be used as crosslinker for hydrogel
formation.[151]
So thiol-norbornene click reaction was applied for forming hydrogels using
different growth factors like TGF-β1 (which induces chondrogensis of hMSCs), matrix
metalloproteinase (MMP) sensitive/cleavable peptides due to the accessibility of thiol groups
to these bioactives.[85, 152]
Figure 14. (a) The reaction mechanism of thiol-Michael addition using base and nucleophile
as a catalyst. Adapted with permission from ref. [153]
. Copyright 2012, Royal Society of
Chemistry, (b) The reaction mechanism of UV assisted thiol-ene reaction. Adapted with
permission from ref. [154]
. Copyright 2010, Royal Society of Chemistry.
1.3.5.1.4 Enzyme Catalyzed
Enzymatic crosslinking has attracted attention due to several advantages like substrate
specificity, good control over reaction rates under mild conditions, and high
biocompatibility.[155]
The most commonly used enzyme is horseradish peroxidase (HRP),
which catalyzes cross-linking through oxidative polymerization of phenol derivatives in the
presence of H2O2. Several natural and synthetic polymers have been functionalized with
25
Introduction
tyramine for crosslinking by HRP in the presence of H2O2.[109, 156]
Heparin-dextran and
dextran-hyaluronic acid hydrogels have been prepared by an enzymatic crosslinking with
HRP and H2O2.[99, 107]
They have also been successfully applied as scaffolds for cartilage
tissue engineering. They have shown promise for matrix production and chondrogenesis. Very
few systems have been exploited for use as injectable scaffolds in cartilage tissue engineering
because of the potential cytotoxic effect from a local high concentration of H2O2.[70]
1.3.5.1.5 Strain Promoted Azide-Alkyne Cycloaddition Reactions
In 2004 Betrozzi et al. first applied a strain promoted azide-alkyne cycloaddition (SPAAC)
reaction in biological medium for staining the cell membrane using strained cyclooctyne[157]
and azide after the development of Cu(I) catalyzed cycloaddition reactions of unsaturated
systems with 1,3-dipoles (Figure 15). The high reactivity of cyclooctynes is due to severe
deformation from the ideal 180°. In 2011 Anseth group applied the same chemistry to create
hydrogels and pattern three-dimensional cell microenvironments.[158]
Recently the Haag group
showed the cyto-compatibility of a SPAAC reaction by encapsulating NIH3T3 cells in the
dPG based degradable microgels.[17]
Among all the crosslinking chemistries discussed above,
SPAAC is advantageous because of its fast reaction kinetics and selectivity towards
substrate.[159-160]
No external stimuli/triggers are required for the initiation of SPAAC
reactions such as UV, photo, pH, base, catalysts (metal salts, oxidizing agents), etc.[161]
Although these properties make them ideal candidates for 3D encapsulation of cells for
cartilage tissue engineering, there have been no reports of a SPAAC reaction being applied for
cartilage tissue engineering to date.
26
Introduction
Figure 15. (a) Cycloaddition with azide (SPAAC). (b) Structures of the most commonly
employed cyclooctynes. Adapted with permission from ref. [160]
. Copyright 2010 WILEY-
VCH Verlag.
1.3.6 Hydrogel Degradation
Hydrogel degradation is influenced by several factors such as the number of degradable
linkages, chemistry used for the synthesis of the linker, presence of cells, and the environment
around the hydrogels.[162]
The processes which reverse the gelation mechanism may lead to
degradation of hydrogels with ionic/physical crosslinks gels.[81]
In other situations, hydrogels
are mainly designed to degrade by hydrolysis, enzyme-mediated processes, or a combination
of both. Enzyme mediated degradation is mainly achieved by introducing small peptide
sequences, which can be recognized by the enzymes present in the host tissue (ECM)[85, 151]
or
the natural polymers known to degrade by certain enzymes/proteases in the crosslinked
structure (Figure 16). In this situation, hydrogel degradation can be initiated by the enzymes
secreted by cells. It can lead to localized gel degradation maintaining the overall integrity of
the hydrogel during degradation.[163]
Two common strategies are mainly applied. One is to use
natural polymers like HA or collagen scaffolds which can be degraded by hyaluronidase or
27
Introduction
collagenase, respectively. The other strategy is to introduce small peptides or amino acid
sequences to the hydrogels during crosslinking which can be degraded or cleaved by MMPs
present in tissues.[152, 164-165]
The diffusion of ECM molecules and proteins are mainly
controlled by the mesh size of the gels.[120]
With increase in the mesh size the diffusion also
increases. The degradation profile should complement the secretion of newly synthesized
ECM. The cell laden hydrogels will dissolve if the degradation occurs too quickly. If the
degradation is too slow, ECM depositions take place in the pericellular region which may
affect the cell function.[81]
Figure 16. A schematic of thiol-ene polymerization reaction between star tetra arm PEG
norbornene and matrix metalloproteinases cleavable peptides (containing dithiol)/PEG dithiol
with variable degradability. Reproduced with permission from ref. [85]
. Copyright 2011
Elsevier Ltd.
28
Introduction
The typical approach for the design of hydrolytically degradable hydrogels involves
incorporation of ester bonds in the crosslinks or the polymer backbone.[144]
Mainly alpha
hydroxyl esters (lactic acid, ɛ-caprolactone, trimethylene carbonate), fumarate, and
polyphosphoesters are introduced in the polymer backbone to introduce degradability in the
hydrogels for cell encapsulation.
Mass erosion of hydrogels during degradation starts when a certain number of linkages are
cleaved due to the hydrolytic or enzymatic degradation of susceptible linkages. Several
characteristic features of the mass loss profile can be seen in a typical degradation study of
hydrogels. First the mass loss (%) increases as the linkages start to cleave. After a certain
amount of time, when all the crosslinks have cleaved, maintaining 3D network is not possible
anymore as all the polymer chains have dissolved.[81]
At this point, the mass loss (%)
increases sharply; this phenomenon is called reverse gelation. Several factors control the
degradation profile of hydrogels. By tuning the chemistry of the linkages,[81]
degradation
kinetics can be tailored, i.e., linkers containing lactic acid degrade faster compared to linkers
with caprolactone units.[166]
Peptides with different reactivities towards the same enzymes
and/or proteases can be prepared by alternating the small amino acid sequences.[149]
By
incorporating such different peptide sequences in hydrogels, the enzymatic degradation can be
also tailored.[149]
Changing the degree of crosslinking can vary degradation time, because
highly crosslinked gels need to cleave more numbers of degrading units.[81]
1.3.7 Requirements for Perfect Tissue Engineering Scaffolds
For production of cartilaginous tissue, a 3D environment is required which is mainly provided
by scaffolds. From the above discussions, the scaffold should have the following
characteristics to become a perfect cartilage tissue engineering scaffold:
First, the scaffold should have a controlled degradation profile and can be transplanted at the
right location (defect site) using minimally invasive technique. Next it should span and fill the
defect site. It should allow the diffusion of nutrients and waste. It should promote cell
viability i.e. the gelation should be cytocomptible to chondrocytes or other regenerative cells.
Finally, it should retain the original phenotype of the cells and help in tissue production. It
should adhere and integrate with surrounding native cartilage and provide sufficient
mechanical integrity to the functional joint to work in vivo.[71]
29
Introduction
1.4 Heparin Mimetic Synthetic Polymers
Glycosaminoglycans (GAGs) are highly negatively charged polysaccharides located on the
cell surfaces produced by mammalian cells. They are mainly bound to proteins or lipids or
free in solution. These GAGs form the glycocalyx (GC) of individual cells. Hyaluronic acid,
dermatan sulfate, chondrotin sulfate, heparin/heparan sulfate, and keratan sulfate are the
relevant GAGs present in mammalian tissue. GAGs have diverse functions such as they
provide structural integrity to cells, help in cell migration, act as lubricating fluids in the joint,
bind and store the proteins with basic amino acids in the extracellular network.[167-169]
Some
GAGs have therapeutic efficacy. For example, heparin has been used as a drug of choice in
the prevention and treatment of thromboembolic disorders.[170]
But the main problem from
using heparin is that it is isolated from animals.[171]
Heparin is not a well-defined biopolymer
and has a range of molecular weights.[171]
Charge and quality also vary from batch to batch.
To overcome those problems, there is an ongoing interest in mimicking the properties of
heparin/heparan sulfate using synthetic polymers.[172-173]
Figure 17. Structure of sulfonated and sulfated polymers used to mimic the properties of
heparin. Adapted with permission from ref. [174]
Mainly the approaches for mimicking heparin have concentrated on using sulfated
glycopolymers, polystyrene sulfonated (PSS) based linear polymers, and sulfonated
30
Introduction
dendrimers.[174]
Recently Varghese et al. prepared PSS based hydrogels and demonstrated the
adhesion of human pluripotent stem cells (hPSCs) and their long-term growth on these
hydrogels.[175]
In recent years, the Haag group has also developed heparin mimetic branched
sulfated polymer dPGS in one step by sulfating dPG (Figure 17).[171]
Until now dPGS is
mainly used as an anti-inflammatory compound and has shown L-and P-selectin
inhibition,[176-177]
inflammation targeting (in an arthritis model),[178]
compliment
activation,[179]
and as well as anticoagulation effects, etc.[180-181]
The main advantage of dPGS
over the other heparin mimetic scaffolds is the easy access to functional groups on the dPG
surface and facile control over the degree of sulfation.[182]
Also its reduced anticoagulate
effect as compared to Heparin is of interest. So far the chemistry that is used to prepare
heparin mimetic hydrogels is unsuitable for 3D cell encapsulation. There is still a need to
develop dPGS based hydrogels using a bio-orthogonal chemistries for a range of biomedical
applications such as tissue engineering, stabilization and delivery of growth factors and
proteins, etc.[183]
This thesis will circumvent these problems and provide a synthetic platform
to access the dPGS based hydrogels for 3D cell encapsulation and will show their efficacy as
a cartilage tissue-engineering scaffold.
31
Scientific Goals
2 Scientific Goals
Hydrogels based on synthetic polymers are interesting for biomedical applications due to their
physical, chemical stability, and versatility in fabrications.[12, 184-186]
Among them, dendritic
polyglycerol (dPG) is a promising candidate due to its high biocompatibility and presence of
multiple hydroxyl groups on the periphery, which can be easily functionalized with different
cationic and anionic functional groups to mimic the properties of natural glycopolymers.[6, 8,
171, 176, 187] The cationic and anionic functional groups can interact with different bio-molecules
which make them attractive candidates for application in tissue engineering, regenerative
medicine, and growth factor delivery, etc.[175]
Above all, dPG is resistant towards plasma
proteins.[7]
The interaction has been studied by immobilizing dPG and oligoglycerols on solid
surfaces.[5, 188]
Recently encapsulation of therapeutically relevant proteins lysozyme and
asparaginase in dPG based nanogels has been shown to keep their original secondary structure
which can lead to protein stabilization.[9]
Thus dPG is an excellent material for protein
(enzyme) immobilization/encapsulation.
Figure 18. Projects within this PhD thesis. Inset: Cells reside in a complex biophysical and
biochemical environment. This microenvironment includes degradable structural fibers,
adhesive binding domains, and proteoglycans for biomolecule sequestration. The picture is
adapted with permission from ref. [184]
. Copyright 2015, Biomedical Engineering Society.
32
Scientific Goals
Considering the benefits of dPG, within the first project, a dPG-based scaffold has been
designed to entrap catalytically active enzymes such as periplasmatic aldehyde
oxidoreductase (PaoABC) and to develop enzyme based biosensors. The main goal is to
provide the enzyme native physiological conditions on an electrode surface. This native
environment can be generated by forming a three-dimensional (3D) network of hydrophilic
polymers. A 3D network was required, as immobilized proteins or enzymes are limited in
monolayers. Therefore, in this project, dPG and PEG (the structures are shown in Figure 16)
will be applied to form a crosslinked 3D network on gold electrodes (surfaces) in order to
entrap PaoABC as a model enzyme. dPG and PEG will be chosen to act respectively as a
multifunctional building block and bifunctional crosslinker. Benzaldehyde could be detected
amperometrically using this immobilized enzyme-polymer construct on an electrode. The
crosslinking approach had to be fast, high yielding, and non-reactive towards the enzymes.
Furthermore, the biosensor response shall be tuned by varying the crosslinking density and
the length of crosslinker, i.e., the length of PEG.
Within the second project, a suitable dPG-based scaffold will be developed using SPAAC for
mimicking the microenvironment around the chondrocytes using synthetic polymers for
degenerative joint diseases. ECM is mainly composed of proteoglycans, collagen, and
bioactive proteins as discussed in Section 1.3. In recent years we developed highly branched
negatively charged polyanion (dPGS) from the sulfation of dPG, which is analogous to
heparin/heparan sulphate proteoglycan. Furthermore, the interaction of human chondrocytes
shall be studied by encapsulating them in the dPGS based hydrogels during gelation. The
amount of dPGS in the hydrogels can affect the tissue formation. So, hydrogels with varying
dPGS content shall be prepared maintaining the overall polymer content constant to
understand the effect of dPGS in terms of tissue formation.
In tissue engineering, the tissue formation can be enhanced by introducing cell-cell and cell-
matrix contact. Cell-cell contact can be incorporated by introducing degradation in the
hydrogels. The main problems associated with current polymerization techniques are the
generation of reactive oxygen species that may lead to cell death (as discussed before in
Section 1.3.5). Thus in the third project, the goal is to develop a universal cyclooctyne
terminated degradable linker that can be used to prepare hydrogels with any azide containing
polymers. The degradable linker will be synthesized by introducing caprolactone blocks in the
PEG chains. Finally the degradation of the dPGS/star PEG hydrogels and cytocompatibility of
the gelation shall be studied.
33
Manuscripts & Publications
3. Manuscripts & Publications
In the following section the published articles and submitted manuscripts are listed and the
contributions of the author are specified.
3.1 Dendritic Polyglycerol−Poly(ethylene glycol)-Based Polymer Networks for
Biosensing Application
Pradip Dey, Miriam Adamovski, Simon Friebe, Artavazd Badalyan, Radu-Cristian Mutihac,
Florian Paulus, Silke Leim hler, Ulla Wollenberger,* and Rainer Haag*
Figure 18. Copyright © 2014, American Chemical Society. Used with permission from ref.
[189].
Author’s contributions: In this publication the author contributed to the concept, and all the
synthesis, characterization, size measurements, electrode modification, the data evaluation, as
well as the draft of the manuscript.
P. Dey, M. Adamovski, S. Friebe, A. Badalyan, R.-C. Mutihac, F. Paulus, S. Leimkühler, U.
Wollenberger, R. Haag, ACS Appl. Mater. Interfaces, 2014, 6, 8937-8941.[189]
http//:dx.doi.org/10.1021/am502018x
34
Manuscripts & Publications
3.2 Mimicking of Chondrocyte Microenvironment Using In Situ Forming Dendritic
Polyglycerol Sulfate Based Synthetic Polyanionic Hydrogels
Pradip Dey,# Tobias Schneider,
# Leonardo Chiappisi, Michael Gradzielski, Gundula Schulze-
Tanzil,* and Rainer Haag*
Figure 19. Structure used from Dey et al.[190]
Author’s contributions: In this publication the author contributed to the concept, and
performed all the syntheses, characterization, the data evaluation, and as well as wrote the
manuscript.
P. Dey, T. Schneider, L. Chiappisi, M. Gradzielski, G. Schulze-Tanzil, R. Haag, Macromol.
Biosci., 2015, In press.[190]
62
Manuscripts & Publications
3.3 Hydrolytically Degradable, Dendritic Polyglycerol Sulfate based Injectable
Hydrogels using Strain Promoted Azide-Alkyne Cycloaddition Reaction
Pradip Dey,* Shabnam Hemmati-Sadeghi, and Rainer Haag*
Figure 20. Structures used from Dey et al [191]
.
Author’s contributions: In this publication the author contributed to the concept and design,
and performed all the synthesis, characterization, the data evaluation, as well as wrote the
draft of the manuscript.
P. Dey,* S. Hemmati-Sadeghi, and R. Haag*, Polym. Chem., 2015, In press.[191]
105
DOI: 10.1039/C5PY01326G
Summary & Conclusion
4. Summary & Conclusion
Considering the bio-inertness of dPG, the main objective of this thesis was to build new
synthetic hydrogels/polymeric scaffolds on the electrode surface for immobilization of
catalytically relevant enzymes. At the same time, introduction of specific functional groups
like sulfate, phosphate, phosphonate, sulfonate, carboxylate, amine, etc. to dPG were
supposed to lead to highly biocompatible structures which would have specific interaction
with biomolecules like siRNA, proteins (containing basic amino acids). Sulfation of dPG led
to formation of dPGS which mimic the properties of heparin/heparan sulfate. Due to such
specific interactions with the proteins, dPGS based hydrogels are interesting for tissue
engineering. So, the aim of this thesis was to develop hydrogels based on dPGS using
bioorthogonal crosslinking chemistry.
In the first project, an easy and fast crosslinking approach was used for the formation of dPG-
PEG based hydrogel film on the gold electrode to entrap catalytically relevant enzyme
periplasmatic aldehyde oxidoreductase (PaoABC) for enzyme based biosensor application.
The 3D polymeric network was used to increase the enzyme loading more than the SAM
approach because the amount of enzyme immobilized in a monolayer is limited. The main
reason for designing a dPG−PEG-based polymer matrix for enzyme encapsulation was that
the enzymes (proteins) do not alter their secondary structure within the polymer matrix due to
less interaction with the matrix and thus lead to maximum protein stabilization. Benzaldehyde
in the concentration range of 0.8−400 μM can be detected amperometrically using this
developed biosensor. By varying all of the parameters like enzyme loading, pH, cross-linking
density, and cross-linker lengths, it was possible to optimize the biosensor performance. The
optimum crosslinking density was found to be 1:6.3 (dPG:PEG ratio). PEG (n = 136) was the
optimum crosslinker for the 3D scaffold formation and it resulted in high amperometric
signals with a wide range of concentrations and very short response times (< 5 s). Using this
dPG-PEG based hydrogel film, the enzymatic activity was retained up to 4 days on the
electrode surface.
In the second project, chondrocyte’s microenvironment was mimic ed using the dPGS based
polyanionic hydrogels. The interaction of this hydrogel with human chondrocytes was also
studied in collaboration with the group of Dr. Gundula Schulze-Tanzil at Charité. In this
project, four different hydrogels with varying amounts of dPGS content ranging from 4 to
35% were developed using bio-orthogonal SPAAC reactions and the PEG hydrogel without
dPGS was used as a control. The mass swelling ratio of the hydrogels also increased with the
128
Summary & Conclusion
incorporation of dPGS in the hydrogels due to increase in the osmotic pressure inside the
hydrogels. By varying the dPGS content, it was possible to vary the elastic moduli of the
hydrogels in the range of 1kPa to 5 kPa. Human chondrocyte encapsulation showed better cell
viability in dPGS containing hydrogels than the pure PEG scaffolds after 21 days in vitro. The
round morphology of chondrocytes can be kept in the hydrogels for 21 days. Furthermore, the
incorporation of dPGS in the hydrogels led to a significant reduction of collagen type I protein
expression. All the hydrogels expressed the collagen type II protein. There was no significant
difference in the collagen II gene expression for all the hydrogels. Collagen II deposition was
mainly observed in the pericellular region in the dPGS incorporated hydrogels due to tight
crosslinking. The higher cell viability in the dPGS hydrogels can be explained by the increase
in swelling behavior or the hydrogels acting as a reservoir for the proteins expressed by the
encapsulated cells. Further studies are needed to understand the dPGS-cell or dPGS-protein
interaction.
In the third project, the main goal was to synthesize a hydrolytically degradable strained
cyclooctyne linker that could be used to react with azide containing polymers for the
preparation of hydrogels, microgels, or nanogels by biorthogonal SPAAC. The degradability
was achieved by introducing polycaprolactone (PCL) units to PEG. The cyclooctynes groups
were introduced by coupling a cyclooctyne-alkyne derivative to azides using a novel
protection-deprotection approach of cyclootynes. At first the cyclooctyne groups were
protected using Cu(I) catalyst, and the excess of Cu(I) was used for the coupling of alkyne
and azide. Afterwards cyclooctynes were regenerated by decomplexing the Cu. Highly
crosslinked structures of the hydrogels were visualized using SEM. Hydrogel degradation was
monitored gravimetrically in 10% FCS. dPGS containing gels degraded at a slower rate
compared to the neutral gels. The cyto-compatibility of the hydrogels was proven by
encapsulating mouse fibroblast L929 cells during the gelation.
129
Outlook
5. Outlook
Polyglycerol scaffolds have proved to be a promising matrix for enzyme immobilization as
demonstrated in the first part of this thesis. Soluble mediators have been used for electron
transfer from the enzyme to the electrode. But it is desirable to have scaffolds containing the
redox species to mediate the electron transfer for applications. So, a future direction will be to
form hydrogel films that contain redox mediators, such as ferrocene, osmium complexes on
the electrode surfaces. The redox mediators can be introduced to the dPG utilizing their
multiple hydroxyl groups.
dPGS hydrogels have shown real promise for further application in tissue engineering (second
and third part of this thesis). There is a need to use the hydrolytically degradable dPGS
hydrogels for the encapsulation of chondrocytes or other regenerative cells such as
mesenchymal stem cells (hMSCs) and evaluate their efficiency as a cartilage tissue
engineering scaffolds in terms of tissue formation. dPGS can easily be uptaken by cells. It can
accumulate at the inflammed site as well as can penetrate the cartilage. Furthermore dPGS can
act as a drug transporter which make dPGS based degradable nanogels interesting candidates
for several biomedical applications. These nanogels can be formed using the same
bioorthogonal chemistry for the encapsulation of therapeutically relevant proteins or
appropriate drugs for the delivery at the inflammed sites.
130
Zusammenfassung
6. Zusammenfassung
Im Rahmen dieser Doktorarbeit wurden Hydrogele basierend auf dendritischem Polyglycerol
(dPG) entwickelt, welche unter Verwendung verschiedener Vernetzungsreaktionen, wie z. B.
durch Amidverknüpfung oder durch ringspannungsvermittelte Azid-Alkin Cycloaddition
(strain promoted azide alkyne cycloaddition, SPAAC), hergestellt wurden. Das
Anwendungsspektrum dieser Hydrogele reicht von der Entwicklung enzymbasierter
Biosensoren bis hin zu Trägermaterialien für Knorpelzellen.
Enzymbasierte Biosensoren wurden für die amperometrische Detektion von Benzaldehyd in
Konzentrationsbereichen von 0.8 bis 400 µM entwickelt. Dafür wurde periplasmatische
Aldehyd Oxidoreduktase (PaoABC) in einem dPG-PEG basiertem Hydrogel Film auf einer
Gold Elektrode eingeschlossen. Optimiert wurde die Leistung dieses Biosensors durch
Variation aller Parameter, wie etwa Enzymbeladung, pH-Wert, Vernetzungsdichte, sowie die
Länge der vernetzenden Einheiten. Als optimale Vernetzungsdichte wurde ein Verhältnis von
1:6.3 (dPG:PEG) ermittelt. Eine optimale Vernetzung wurde mit PEG mit n=136 erreicht,
wobei amperometrische Signale über einen weiten Konzentrationsbereich und gleichzeitig
sehr kurzen Antwortzeiten (<5 s) beobachtet werden konnten.
Im zweiten Teil dieser Doktorarbeit wurden polyanionische Hydrogele basierend auf
variierenden Anteilen dendritischer Polyglycerinsulfate (dPGS) entwickelt, um unter
Verwendung bioorthogonaler Reaktionen die natürliche Umgebung der Chondrozyten zu
imitieren. Die Bildung des Hydrogels wurde mittels oszillatorischer Rheologie kontrolliert,
wobei die Gelbildung innerhalb von 5 bis 10 min erfolgte. Durch Veränderung des dPGS
Anteils konnten die Werte des Elastizitätsmoduls in einem Bereich von 1-5 kPa variiert
werden. Die Effektivität der dPGS basierten Hydrogele wurden als Knorpel für das Tissue
Engineering evaluiert, indem humane Chondrozyten während der Gelbildung verkapselt
wurden. Hierbei behielten die Chondrozyten in allen Hydrogelen im Verlauf von 21 Tagen in
vitro ihre Viabilität sowie ihre ursprünglich runde Morphologie. Im Vergleich zu den
Kontrollexperimenten (z.B. rein PEG-basierte Hydrogele, Alginat-basierte Hydrogele) wiesen
die Hydrogele mit inkorporiertem dPGS nach 21 Tagen die höchste Zellviabilität auf. Des
Weiteren wurden die Proteinexpression von Kollagen Typ I und Typ II sowie die gesamte
Kollagenexpression der Chondrozyten bestimmt. Die Expression von Kollagen Typ I war mit
dPGS inkorporierenden Hydrogelen signifikant reduziert, d.h. dPGS Hydrogele waren in der
Lage die Dedifferenzierung der Chondrozyten zu inhibieren. dPGS basierte Hydrogele
erschienen vielversprechend als Ger st f r das “Tissue Engineering” der Knorpelzelle.
131
Zusammenfassung
Durch Einführung von ɛ-Caprolacton-Einheiten in PEG sind die dPGS-Hydrogele abbaubar,
außerdem erhöhen die Caprolacton-Einheiten den Zellkontakt und verbessern die
Gewebebildung. Hierfür wurden Linker aus PEG-Polycaprolacton, terminiert mit gespanntem
Cyclooctin (PEG-PCL-DIC) synthetisiert, wobei die Cyclooctingruppe durch Anwendung
einer Schutzgruppenstrategie des gespannten Cyclooctins eingeführt wurde. Die
Abbaubarkeitsstudie in vitro zeigte, dass die mit dPGS inkorporierenden Hydrogele im
Vergleich zu PEG Hydrogelen langsamer abgebaut wurden. Die Zytokompatibilität der
Hydrogelbildung wurde durch Verkapselung von Maus Fibroblasten der Linie L929
nachgewiesen.
Zusammenfassend wurden in dieser Doktorarbeit Methoden der bioorthogonalen Vernetzung
zur Bildung von dPG-basierten Hydrogelen unter physiologischen Bedingungen gezeigt.
Weiterhin wurde die Synthese eines neuen abbaubaren Linkers mit einem gespannten
Cyclooctin als Endgruppe dargestellt. Dieser kann für die Vernetzung eines jeden Azid-
haltigen Polymers für die Bildung abbaubarer Hydrogele genutzt werden. Vor allem aber
wurde ein neues Cyclooctin-Alkin Derivat entwickelt, welches für die Kupplungsreaktion
zweier Azide genutzt werden kann (Kupplungsreaktionen zweier ähnlicher funktioneller
Gruppen). Diese Chemie kann für die Konjugation zweier Polymere, Polymer-Farbstoff,
Polymer-Wirkstoff, Polymer-Protein, Polymer-Zucker usw. verwendet werden.
132
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Appendix
8. Appendix
8.1 Publications & Manuscripts
(1)
(2)
P. Dey, M. Adamovski, S. Friebe, A. Badalyan, R. Mutihac, F. Paulus, S. Leimkühler,
U. Wollenberger, and R. Haag, Dendritic Polyglycerol−Poly(ethylene glycol)-Based
Polymer Networks for Biosensing Application, ACS Appl. Mater. Interfaces, 2014, 6,
8937-8941. DOI: 10.1021/am502018x
P. Dey,# T. Schneider,
# L. Chiappisi, M. Gradzielski, G. Schulze-Tanzil,* and R.
Haag*, Mimicking of Chondrocyte Microenvironment Using In Situ Forming
Dendritic Polyglycerol Sulfate Based Synthetic Polyanionic Hydrogels, Macromol.
Biosci., 2015, In press. # - equally contributed.
(3) P. Dey,* S. Hemmati-Sadeghi, and R. Haag*, Hydolytically Degradable, Dendritic
Polyglycerol Sulfate based Injectable Hydrogels using Strain Promoted Azide-Alkyne
Cycloaddition Reaction, Polym. Chem., 2015, In press. * - Corresponding authors.
8.2 Conference Contributions
Oral Presentations
(1) International Dendrimer Symposium 2015 (IDS-9), Montreal, Canada (12-17 July),
Dendritic Polyglycerolsulfate (dPGS) based Injectable Hydrogels for Cartilage Tissue
Engineering; P. Dey, T. Schneider, L. Chiappisi, M. Gradzielski, G. Schulze-Tanzil,
and R. Haag
(2) European Polymer Federation (EPF) 2015, Dresden, Germany (21-26 June), Heparin
mimetic dendritic polyglycerol sulfate based injectable hydrogels for cartilage tissue
engineering; P. Dey, T. Schneider, L. Chiappisi, M. Gradzielski, G. Schulze-Tanzil,
and R. Haag
(3) Euro BioMAT 2015: European symposium and exhibition on biomaterials and related
areas, Weimar, Germany (21-22 April), Heparin mimetic dendritic polyglycerol
sulfate based injectable hydrogels for cartilage tissue engineering; P. Dey, T.
Schneider, L. Chiappisi, M. Gradzielski, G. Schulze-Tanzil, and R. Haag
Poster Presentations
(4) Biochemie 2014: Bioorthogonal chemistry, Berlin, Germany (16-18 July),
Bioorthogonal Synthesis of Dendritic Polyglycerolsulfate (dPGS) based Hydrogels; P.
Dey, T. Schneider, L. Chiappisi, M. Gradzielski, G. Schulze-Tanzil, and R. Haag
(5) 4th BSRT PhD symposium: Regeneration is Communication: Fireside Chats between
Cells & Matrices, Berlin, Germany (04-06 December 2013), Dendritic
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Appendix
Polyglycerolsulfate (dPGS) based Hydrogels for Cartilage Tissue Engineering; P.
Dey, T. Schneider, L. Chiappisi, M. Gradzielski, G. Schulze-Tanzil, and R. Haag
(6) Engineering Life 2013: Bio-molecular principles for novel methods and materials,
Dresden, Germany (17-18 September, 2013), Dendritic Polyglycerolsulfate Hydrogels
for Cartilage Tissue Engineering; P. Dey, T. Schneider, L. Chiappisi, M. Gradzielski,
G. Schulze-Tanzil, and R. Haag
(7) Tag der Chemie-2013; Universität Potsdam, Germany (27 June 2013), Polyglycerol
Scaffolds for the Immobilization of Catalytically Active Enzymes; Pradip Dey, Miriam
Adamvoski, Ulla Wollenberger, and Rainer Haag
(8) ISHHC XV - International Symposium on the Relations between Heterogeneous and
Homogeneous Catalysis XV, Berlin, Germany (11-16 September 2011) Development
of biosensors from the immobilization of enzyme in polyglycerol scaffolds; Pradip Dey,
Dirk Steinhilber, Ulla Wollenberger, and Rainer Haag
(9) Joint International Symposium - CRC 546 - CoE UniCat Activation of Small
Molecules - Gas Phase Clusters, Molecular Catalysts, Enzymes and Solid Materials;
Erkner, Germany (20-23 February 2011), Novel polymeric supports from the
immobilization and stabilization of chemo- and biocatalysts as well as cells for
catalytic reactions
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Appendix
8.3 Curriculum-Vitae
For reasons of data protection, the curriculum vitae is not included in the online version
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